The present invention relates to the magnetic resonance arts. It finds particular application in conjunction with medical diagnostic magnetic resonance imaging and will be described with particular reference thereto. However, it is to be appreciated that the present invention also finds application in magnetic resonance spectroscopy and magnetic resonance imaging for other applications.
Heretofore, magnetic resonance imagers included a series of annular magnets which generated a temporally constant, generally uniform magnetic field through their common bores. In superconducting magnet designs, the magnets were encased in a vacuum dewar. Whether in a vacuum dewar or other housing, a patient receiving bore was defined through the magnets. In general, the longer the bore, the more uniform the magnetic field. Typically, the bore length was on the order of 1.5 to 2.0 meters. The uniformity of the magnetic field is generally characterized in terms of the amplitudes of zonal harmonic field coefficients Z.sup.0, Z.sup.1, Z.sup.2, normalized to a defined diameter spherical volume.
Even such "long bore" magnets had some magnetic field inhomogeneities. Some of the inhomogeneities were attributable to build or manufacturing errors, and some were due to limiting or theoretical characteristics of the magnet design. Such long bore magnets typically had harmonic distortions, due to build errors, on the order of Z.sup.0 through Z.sup.6. To correct these inhomogeneities, ferrous shims were mounted along the patient receiving bore. These lower order harmonics were cured with a minimal amount of ferrous material. In some imagers, the shims were mounted inside the bore and in others they were mounted outside the bore, e.g., in a superconducting magnetic's cryo-vessel. Typically, shim trays were constructed of a non-ferrous material. The pockets of the non-ferrous trays received ferrous shims, e.g., thin plates of steel coated with a non-conductive polymer or oxide coating. The shims were constrained in the pockets and the trays were mounted inside or outside of the bore.
In magnetic resonance imagers, a magnetic field gradient coil was mounted inside the magnet's bore. Typically, the gradient coil assembly was mounted radially inward from the shims. Imagers with self-shielded gradient coils had a main or primary magnetic field gradient coil and a shield or secondary magnetic field gradient coil which were mounted in a spaced relationship. The shield gradient coil was often placed outside of the bore and, in a superconducting magnet imager, could be placed within the cryo-vessel. With self-shielded gradient coils, the shim trays were often positioned between the primary and shield gradient coils. See, for example, U.S. Pat. No. 5,349,297.
A radio frequency coil was mounted radially inward from the gradient coil(s). A radio frequency shield was mounted between the radio frequency coil and the gradient coils. The radio frequency shield blocked radio frequency signals from reaching the gradient coils, the shims, and other surrounding constructions in which eddy currents could be induced. Radio frequency eddy currents would generate radio frequency signals which would be transmitted into the interior of the radio frequency coil. The eddy radio frequency signals lowered the signal-to-noise ratio during reception of resonance signals and increased the power demands on the radio frequency coil during radio frequency transmission. In addition, the radio frequency shield blocked the transmission of any noise conducted into the bore via the gradient coils.
A cosmetic liner was commonly positioned radially inside the radio frequency coil to prevent the imaged patient from touching the radio frequency and gradient coil constructions and for cosmetic purposes.
One of the difficulties with such magnetic resonance imagers is that the 1.5 to 2.0 meter long bores were claustrophobic to many patients. The long bores also prevented medical personnel from accessing the patient while in the bore. Typically, to perform a medical procedure based on the diagnostic images, the patient needed to be removed from the bore and the diagnostic images reregistered with the patient in the new patient position. If a probe, such as a biopsy needle, was inserted into the patient and the physician wanted to check with the magnetic resonance imager whether it was inserted to the proper location, the patient and probe needed to be reinserted into the bore for another imaging session.
One solution to these problems resides in the use of "short bore" magnets, e.g., 1.25 meters or less. Although short bore magnets render the imager more user friendly and provide improved access to the patient, the temporally constant magnetic field generated by the short bore magnets tends to be less homogeneous. Moreover, such short bore magnets typically have higher than Z.sup.6 order harmonic distortions, intrinsic to the magnet design, to be shimmed. Higher order distortions require significantly more ferrous material for shimming than do Z.sup.6 and lower order distortions.
The present invention overcomes the above-referenced problems and others.